This invention relates to a method and apparatus for manufacturing superconducting magnet coils for use in magnetic resonance imaging and, more particularly, to a method for overwrapping an embedded b-zero coil about the superconducting magnet coil.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field b-zero, b0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but process about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. An NMR signal is emitted by the excited spins after the excitation signal B1 is terminated and, this signal may be received and processed to form an image.
As is well know, a magnet can be made superconducting by placing superconducting magnet coils in an extremely cold environment, such as by enclosing it in a cryostat, or pressure vessel, containing liquid helium or other cryogen. The extreme cold reduces the resistance in the magnet coils to negligible levels, such that when a power source is initially connected to the coil (for a period, for example, of ten minutes) to introduce a current flow through the coils, the current will continue through the coils due to the negligible resistance, even after power is removed, thereby maintaining a magnetic field. Superconducting magnets find wide application, for example, in the field of medical magnetic resonance imaging (hereinafter called “MRI”).
As is also well known, MRI requires very strong or large magnetic fields in the imaging bore. Such fields require a very high degree of uniformity or homogeneity. Such fields are, however, also subject to the effects of external disturbances. The superconducting magnetic field is generated by current flow through superconducting magnet coils, each having as many as 10,000 turns. The many turns and layers of turns must be insulated from each other and the insulation integrity must be maintained during superconducting operation, including the ramping up of the coils to the operating current, and the possible sudden quenching, or discontinuance, of superconducting operation. Any shorting of magnet coil turns can produce heat which can lead to very serious inadvertent quenching of magnet operation which leads to the sudden escape of helium and the need to replenish the liquid helium and again ramp the MRI magnet up to operating current entailing significant equipment down time and expense. Such shorting can also damage the superconducting wire, rendering the MRI magnet inoperable. Because of the extreme high current levels involved in the high strength magnetic fields, the superconducting magnet coils must undergo extreme and difficult operating environment and forces, including significant thermal, magnetic, electrical and mechanical forces generated during such operations. As a result, the cost of producing superconducting magnetic coils and magnets is relatively high, frequently costing into the hundreds of thousands of dollars for a single MRI magnet assembly.
The homogeneity requirement for the magnetic field of the MRI cannot be achieved simply by controlling manufacturing tolerances. In practice, extra coils, typically called correction coils and/or passive shims, are provided to correct or improve the magnetic homogeneity. This allows for reasonable manufacturing tolerances. Additionally, b-zero coils are used on superconducting MRI magnets to minimize the effect of external disturbances. The effectiveness of the superconducting b-zero coil is dependent on the location in relation to the main field windings of the coil. In methods of current production, b-zero coils are wound on a flat surface and a fiberglass structure is formed over the coils. This method limits the ability of a designer to optimize the location of the b-zero coils. Accordingly, the limitations on the b-zero coil location limit the ability of the b-zero coil to effectively reduce external disturbances to the magnetic field.
Furthermore, considerable engineering and development has been applied for some time to improving and simplifying, to the extent possible, the magnetic systems for MRI devices so as to minimize the effects of external disturbances in order to improve imaging quality with uncomplex means and without additional expense. In the method of current magnet design, such analysis shows that one of the most effective locations for the b-zero coil is directly on top of the bucking coil superconducting wire. Unfortunately, a layer of aluminum overwrap is required on this surface to provide structural integrity to the superconducting coil. Placing the thin b-zero wire, which is typically manufactured as a thin copper wire having a Formvar® or similar insultation, between the main superconducting coil and the aluminum overwrap would most likely result in damage to the wire due to the highly compressive forces that develop during coil winding, cool down and magnet operation as previously described.
What is needed is a method for winding an embedded b-zero coil about a main coil that maintains the integrity of the superconducting coil and the b-zero wire during coil winding and also during normal operation of the superconducting MRI magnet. What is also needed is such a method that is not location sensitive but that allows for placement of the b-zero wire wherever such is desired or required by the designer, and at a cost savings to overall production.